1. Field of the Invention
The invention relates to the field of sensors. In particular the invention relates to methods and apparatus for sensing biological and chemical species, as well as the measurement of viscosity.
2. Description of the Related Technology
Most of the current work done with biosensing technologies relies on fluorescence, lasers, fiber-optics-based methods, quartz crystal microbalance technology, electrochemical enzyme immunoassays, and binding to metal particles. Most of these techniques are neither direct, nor quantitative. Many of these techniques are also quite slow. In addition, most of the aforementioned techniques do not lend themselves to measurement of changes in mass, which may provide a convenient way to measure a variety of different parameters.
A mass sensor based on resonance frequency needs three components, an actuator (driver), a resonator, and a detector. A popular mass sensor is a silicon-based micro-cantilever due to its commercial availability and ease of integration with existing silicon based methodologies. In a silicon-based micro-cantilever mass sensor, the micro-cantilever acts as the resonator and is driven by an external lead zirconate titanate (PZT) actuator at the base of the micro-cantilever to generate vibrations in the resonator, which may be detected by an external optical detector. For bio-detection, receptors are immobilized at the micro-cantilever surface. Binding of antigens to the receptors immobilized on the cantilever surface increases the cantilever mass and causes a decrease in the resonance frequency. Detection of target molecules is achieved by monitoring the mechanical resonance frequency. In spite of the popularity of silicon-based micro-cantilevers, they rely on complex external optical components for detection. In addition, the PZT vibration driver adds to the weight and complexity of the sensor. Further, the external actuator can only be located at the base of the micro-cantilever, which greatly limits its effectiveness in driving the cantilever's vibration. The optical means of detection also limits how small the micro-cantilever can be fabricated, and therefore limits the mass detection sensitivity.
In addition to mass detection, silicon-based micro-cantilevers have also been used as sensors for small molecules by detecting the stress generated on the cantilever by the adsorption of species onto receptors associated with the cantilever. Antibody or DNA receptors are coated on the surface of the micro-cantilevers to bind target protein or DNA molecules. The stress generated at the time of binding or unbinding of the target molecules to the receptors on the micro-cantilever surface induces a temporary deflection of the micro-cantilever that may be detected by external optical components or by an adsorption-stress-induced DC voltage on a piezo-resistive coating layer on the cantilever surface. Because the binding-induced stress decays with time, it can only be detected when the micro-cantilever is first introduced to the target molecules. The induced stress, and hence the induced DC voltage, dissipates within 20 minutes. Also, detecting the adsorption-induced stress in this manner offers no information about the amount of target antigen adsorbed on the cantilever.
Moreover, immersing silicon-based micro-cantilevers in water reduces the resonance intensity by an order of magnitude, reducing the Q factor, defined as the ratio of the resonance peak frequency relative to the resonance peak width at half peak height, to about one, thus making it impractical to use silicon-based micro-cantilevers for in-water detection. The main reason that such silicon-based micro-cantilevers do not exhibit sufficiently high resonance signals in water is that silicon-based micro-cantilevers are not piezoelectric. The deflection at the tip of the silicon-based cantilever is driven by the vibration driver located at the base of the silicon-based cantilever and is detected by external optical components. Driving a cantilever at its base is not the most effective way to generate deflections at the tip of the cantilever. While the relatively weak deflection signal generated by the vibration driver at the base is sufficient for in-air detection it does not withstand the damping of water.
Silicon-based micro-cantilever sensors therefore have the following shortcomings when used for mass detection using the resonance mechanism: (1) A silicon-based micro-cantilever sensor needs to be driven by an external actuator. (2) A silicon-based micro-cantilever sensor loses its detection sensitivity in water due to viscous damping. (3) A silicon-based micro-cantilever requires a complex external optical detection system. The resolution of the optical detector puts a limit on how small the displacement can be and therefore, how small the cantilever can be, which places significant constraints on detection sensitivity.
For detection of stress, the silicon-based micro-cantilever sensor does not need a driver, but requires an external optical system or a piezo-resistive layer. Stress detection using a piezo-resistive layer involves DC electrical measurements. Furthermore, detection using a piezo-resistive layer is not very sensitive. Therefore, most silicon cantilevers use an external optical means for detection. In silicon-based micro-cantilevers, the adsorption induced stress decays in 20 minutes and the adsorbed amount cannot be quantified.
Compared to silicon-based sensors, piezoelectric micro-cantilever sensors are not as bulky and complex. Piezoelectric devices are excellent transduction candidates because of their short response time and high piezoelectric coefficients. Because they are piezoelectric, both the driving and sensing of the mechanical resonance can be conveniently done electrically within the resonator. Currently, piezoelectric biosensors are based on commercially available quartz crystal microbalances (QCM), a disk device that uses thickness-mode resonance for sensing. Although quartz is a weak piezoelectric material, it is widely used as a layer thickness monitor in part due to the availability of large quartz single crystals to make the membranes. The typical mass detection sensitivity of a 5 MHz QCM that has a minimum detectable mass density (DMD) of 10−9 g/cm2 is about 10−8 g/Hz, about two orders of magnitude less sensitive than millimeter sized piezoelectric cantilevers. Moreover, because QCMs are larger in size, they are harder to develop into array sensors for multiple antigens. Quartz is a weak piezoelectric, much like the silicon-based cantilever, when immersed in water, the resonance peak intensity of QCM is reduced to less than one twentieth of the in-air peak intensity due to viscous damping of water, thus limiting the use of QCMs in water.
QCMs employ a piezoelectric crystal that serves as the actuator, resonator, and detector. However, it involves shear waves of the thickness mode of the crystal rather than the flexural mode of the cantilever geometry. QCMs have a lower mass sensitivity than a silicon-based micro-cantilever. QCMs are also limited by the following shortcomings. Due to its planar geometry, for mass detection, QCMs are not capable of detecting very small amounts of mass, which limits the detection sensitivity. Also, QCMs use higher resonance frequencies (>5 MHz), which reduces the relative sensitivity, (Δf/f), where f and Δf denote the initial resonance frequency and the resonance shift. QCMs are also difficult to miniaturize in order to improve the detection sensitivity.
Comparing the two direct biosensor technologies, silicon-based-micro-cantilever sensors1, exhibit high mass-detection sensitivity, but the optical detection system is large and complex. QCM-based sensors have the merit of simple electrical driving and electrical detection but they exhibit a much lower mass sensitivity than silicon-based micro-cantilevers.
QCMs are also larger in size thus harder to develop into array sensors for multiple antigens.
Therefore, there exists a need for improvement of the sensing capabilities of existing sensors.